The present invention relates generally to devices, systems and methods for assisting blood flow, and, particularly, to blood pumps that include a moving (for example, an oscillating) valve assembly to propel blood. The pumps of the present invention can be either fully implanted or temporarily connected to the circulation using percutaneous blood conduits. The pumps of the present invention can, for example, be fully or completely implanted for months to years to alleviate or correct heart failure and related symptoms.
Heart failure, or the inability of the heart to pump sufficient blood for the body's needs, results in very poor quality of life, huge costs to society, and hundreds of thousands of yearly deaths. Heart failure is caused by an abnormally low cardiac output. Cardiac output is the out flow of blood from the heart and is measured in liters of blood flow per minute or I/min. Cardiac output for a normal man at rest or during light activity is around 5 liters per minute. Severe heart failure exists when the cardiac output is as low as 2.5 to 3.5 liters per minute. For an average man in heart failure with a heart rate of 80 beats per minute, the average amount of blood that is pumped with each heartbeat or stroke volume might be 37 milliliters or ml. The same man with a normal heart might pump 62 milliliters with each heartbeat. An ideal treatment for heart failure would boost the low 37 ml stroke volume up to the normal 62 ml stroke volume.
The main pumping chamber of the heart or left ventricle, LV has an inlet mitral valve and an outlet aortic valve. During left ventricular contraction or systole, the inlet valve closes as blood is pushed through the aortic valve into the aorta or main artery to the body. When the LV is resting during diastole, LV pressure may be between 2 and 20 mm of Hg pressure. This diastolic pressure is termed the LV preload and the preload will be in the higher end of its pressure range during heart failure. During active LV contraction or systole, the LV must eject its blood against the pressure in the aorta. Aortic pressure is typically between 70 and 140 mm Hg Pressure. This aortic pressure is termed the after-load. It is well known that, if the after-load is reduced in heart failure, the LV stroke volume will naturally increase and this increase is one reason that afterload-reducing drugs such as ACE-inhibitors help heart failure patients.
Blood pumps which lower the aortic pressure after-load are attractive because they allow the failing LV to eject more blood with less effort. However, no after-load reducing devices have thus far been shown to be practical for indefinitely supporting the failing LV. Instead, all long term (that is, months to years), commercially available heart assist devices, whether rotary turbine pumps or collapsing chamber pumps go around or bypass the failing LV, pumping blood from the LV apex through the pump into the aorta. By doing so, they act in parallel to the LV and essentially compete with the LV in their pumping action. This pumping competition has several negative complications including right heart failure, fusion of the aortic valve over time and the risk of collapsing the LV. Collapsing chamber pumps are physically large and cannot be implanted in some small patient because of their size. Rotary turbine pumps are attractively small, but have other limiting complications. For example, the rotary turbine pumps induce high levels of shear stress in the blood elements and also may reduce the normal pulsatility of the blood entering the aorta. The effect of the high shear stress on the blood cells is to promote blood clotting which can lead to strokes and heart attacks. Physicians try to reduce this blood clotting by giving the patients anticoagulants and this, in turn, puts the patients at risk of excessive bleeding. These clotting and bleeding complications are substantial limitations to broader use of rotary turbine assist pumps.
For short-term heart assist (that is, hours to days), a common method of providing cardiac assist is the use of counterpulsation devices such as intraaortic balloon pumps or IABPs. IABPs provide an afterload-reducing type of assist. As described in U.S. Pat. Nos. 4,733,652 and 3,692,018 by Kantrowitz et al. and Goetz et al., the main benefit of such devices stems from after-load reduction of the left ventricle during systole and providing increased diastolic pressure for perfusing the coronary and other arteries during diastole. Typical patients needing this type of treatment suffer from cardiogenic shock or need perioperative circulatory support. The nature of IABP design restricts itself to acute use only, since the bulky balloon drive mechanism remains outside the patient's body necessitating patient confinement to a hospital bed.
A “dynamic aortic patch” is disclosed in U.S. Pat. No. 4,051,840, to Kantrowitz et al. and is in clinical trials. It is surgically and permanently attached to the patients descending aorta and is pneumatically activated by an external air pump. Such a pump lowers the LV after-load, facilitating left ventricular contraction and increasing stroke volume.
Pouch-type auxiliary ventricles attached to the patient's aorta have been described. These devices use mechanical or pneumatic means for the pumping the blood contained in the pouch and are disclosed in U.S. Pat. Nos. 3,553,736 and 4,034,742 by Kantrowitz et. al. and Thoma. Some of these devices have a single access port to the aorta that serves as both the inlet and the outlet for blood flow. Single port designs have the disadvantage of recirculation and relative flow stagnation, increasing the risk of clot formation and thromboembolism. Others have both inlet and outlet ports to the aorta and are typically connected in parallel with the aorta. See, for example, U.S. Pat. Nos. 4,195,623 and 4,245,622 by Zeff et al. and Hutchins et al.
U.S. Pat. Nos. 5,676,162, 5,676,651, and 5,722,930, by Larson et al., describe a single stroke moving valve pump designed for ascending aortic placement. The Larson device uses a commercially available artificial heart valve with attached magnets and requires excision of a portion of the aorta. A series of separate electric coils step the valve/magnet combination forward in a sliding action within a cylinder. The device is quite large for the limited space available between the heart and the take-off vessels from the aorta to the upper body and brain. The device is designed to have one stroke in synchronization with each LV systole. The blood volume required for closing commercially available heart valves is typically 2-5 ml and therefore multiple smaller oscillations per heart contraction would suffer from volumetric inefficiency. Another problem with the Larson device is the tight crevice between the cylinder wall and the moving valve. This tight space results in high blood shear and the resultant risk of blood clotting complications. The same problem exists with a moving valve pump described by Child, U.S. Pat. No. 4,210,409. The Child pump has two valves, one stationary and one moving.
Thornton, U.S. Pat. No. 5,147,281 discloses an oscillatory valve blood pump that is external to the body and fits in an enclosure the size of a briefcase. It uses a stationary coil to attract a magnetic tube encasing a one-way valve. Its forward stroke propels blood until the tube assembly stops and is repelled backward by return leaf springs that were charged during the forward stroke. A second stationary valve is sometimes in the circuit. A stretchable silicone rubber tube connects the tube or pipe-valve assembly with the pumps inlet and outlet.
Nitta, in ASAIO Transactions 1991:37: M240-M241 describes a “univalved artificial heart” powered electro-magnetically wherein the valve oscillates within the frequency range of 1 to 30 Hz. The valve is contained in a tube, with attached magnetic material. Stationary electric coils actuate the tube-magnet-valve combination. The valve is described as a jellyfish valve. One problem with jellyfish valves is the compound curvature or wrinkling of the membrane that occurs when the valve opens and closes. One can liken the action of the jellyfish valve to that of an umbrella that oscillates between a circular flat membrane and a wrinkled umbrella shape as it closes and opens. Wrinkling of the membrane is virtually impossible to prevent in a jellyfish valve and introduces stresses and strains that significantly limit the life of the valve.
Hashimoto, U.S. Pat. No. 5,266,012, also uses a jellyfish valve in a vibrating pipe blood pump intended for use outside the body. The purpose of this invention is to make the vibrating tube pump portion separable from the drive mechanism so that the blood-contacting portion of the pump can be disposable.
Although numerous pharmacologic, biologic, and mechanical interventions have been devised to address heart disease/failure (some of which are described above), heart failure remains a major public health problem with an estimated five million victims in the United States alone. It is, therefore, very desirable to develop improved devices, systems and methods of assisting the heart in pumping blood through the circulatory system.